Measurement of volumetric fluid flow and its velocity profile

ABSTRACT

A method to measure spatial fluid flow components and their velocity profiles in a number of locations in a cross-sectional area of a lumen or other body cavity by using ultrasound in which the cross-sectional area is interrogated by a plurality of ultrasonic beams; the estimation of the spatial flow components is obtained from a combination of estimations of axial, lateral and total flow; the estimation of one or more flow components is obtained through any combination of time-shift and decorrelation analysis of two or more beam-signals of the interrogating ultrasound transducer; and the estimation accuracy is further improved by the use of a reference decorrelation curve obtained from experiments or beam-theory or both.

[0001] The assessment of fluid flow in the human body is important formedical diagnosis. For example, blood velocity and the volumetric flow(i.e., the volume of blood flowing through a blood vessel, e.g., inliter per second) routinely assist clinical decisions.

[0002] Various ultrasound techniques can be used to measure the motionof scattering materials such as blood, body fluids and tissue.Ultrasound contrast agents can also be used to enhance signals fromfluids with insufficient scatter properties. For example, blood velocitycan be measured in a small volume using the Doppler principle. Inechographic B-scanning, multiple estimates of blood velocity in theplane of the scan can be combined with the gray-scale echo image bycolouring.

[0003] Miniaturised ultrasound transducers can be placed inside thelumen of a vessel or other body cavities to obtain a cross-sectionalecho image. The same ultrasound echo signals can be used to measure thevelocity of the flowing blood or other fluids.

[0004] The purpose of the invention is to provide a method for measuringvolumetric fluid flow and its velocity profile in a lumen or other bodycavity. According to the invention an ultrasonic method is provided tomeasure volumetric flow through a lumen by accomplishing simultaneouslyand in situ (in place) the steps of

[0005] a) measuring the local velocity of the scattering mediumperpendicular to the ultrasound scan plane and

[0006] b) integrating such velocity measurements over the area of thelumen.

[0007] An approximation to the above method in accordance with theinvention comprises the steps of

[0008] a) calculating the average value of the velocity of thescattering medium perpendicular to the ultrasound scan plane,

[0009] b) calculation of the area of flow, and

[0010] c) multiplying the average velocity by the area to obtain volumeflow.

[0011] Furthermore, in order to reduce the number of calculations, theaverage velocity can be approximated by measurement of the velocity in asub-region of the area of flow smaller than the total area of flow, andvolume flow can be computed as before.

[0012] In the method according to the invention the scattering fluid ofinterest may be blood. For purposes of explaining the invention, theinvention shall be discussed in relation to blood. Other fluids may bemeasured in a similar manner. Blood is composed by red blood cells(RBCs), white blood cells and platelets suspended in a liquid calledplasma. Because the size and density of RBCs is large compared to thatof white cells and platelets, backscatter of blood is attributed to thered blood cells. The measurement of blood velocity comprises the stepsof:

[0013] a) obtain (transmit pulse and receive echo) two or moresubsequent echo signals from a single (or a slightly changed) positionof the ultrasound transducer at controlled interval(s) of time Δt,

[0014] b) measure one or more displacements of the blood relative to thebeam, Δd, and

[0015] c) compute velocity from the ratios of displacements over thetime interval, v=Δd/Δt.

[0016] One embodiment of the invention relates to the measurement ofvolumetric flow and velocity imaging from within the lumen of a bloodvessels using intravascular ultrasound. It has to be mentioned here thatconventional ultrasound has already been proposed to measure bloodvelocity on the plane of the scan. The scan is usually oriented alongthe blood vessel. If the scan plane is oriented perpendicular to theblood vessel, volumetric flow could also be computed as describehereinafter.

[0017] The invention will be described in details with reference to theaccompanying drawings, in which

[0018]FIG. 1 is a schematic view of a lumen in which an ultrasoundcatheter is positioned;

[0019]FIG. 2a-2 c are graphs of echo-signals received by an ultrasoundtransducer and of the time shift between echo-pairs;

[0020]FIG. 3 is an illustration of the characteristic decorrelation fora single range;

[0021]FIG. 4 is an illustration of the decorrelation “calibrated”velocity estimation procedure;

[0022]FIG. 5 is a schematic view of a lumen and the scan plane in suchlumen;

[0023]FIG. 6 illustrates the direction of flow relative to thethree-dimensional orthogonal axes centered on the transducer aperture;

[0024]FIG. 7 is a graph showing velocity profiles at different flowvelocities;

[0025]FIG. 8 shows different flow velocity images computed within oneheart cycle; and

[0026]FIG. 9 is a graph showing, simultaneous calculation of phasiccross-sectional area and phasic volumetric flow.

[0027] Local estimates of blood velocity are obtained by means ofecho-signal decorrelation and time-shift analysis. By means of anultrasound transducer, a sound pulse is transmitted into the scatteringmedium; backscattered echoes from the medium are received by the same(or a separate transducer) and converted to an electric signal suitablefor storage and processing. In FIG. 1 a rotational scan of the beam of atransducer is depicted.

[0028] The velocity of a moving object can be calculated by measuringthe displacement of the object during a given interval of time. Theratio of displacement and time interval is the velocity.

[0029] The displacement of an ultrasound-scattering material (such asblood) moving through the beam of an ultrasound transducer results inconcomitant changes in the received echo signal. For example, FIG. 2ashows a sequence of five echo signals (S1 through S5) obtained in anexperiment where a scattering material is progressively displacedthrough an ultrasound beam. It can be observed that with increasingdisplacement, the echo signals progressively a) shift (advance) in time;and b) change in shape (FIG. 1a).

[0030] The correlation coefficient ρ is a measure of the similarity (ordissimilarity) between a pair of signals and is defined such that ρ=1(100%) when there is total similarity, and ρ=0 when the signals are noterelated at all. A decrease in correlation is termed decorrelation.

[0031] In an experimental example, the four correlation coefficients (orsimply correlation for short) computed between echo-signal pairs S1-S2,S1-S3, S1-S4, and S1-S5 are shown in FIG. 2b. Correspondingly, theprogressive time shift between the same echo signal pairs is computedand shown in FIG. 2c, In this experiment the time interval was Δt=250μs. In general the time interval must be sufficiently short to warrantrecognition of the advance in time and shape change of the echo signals:that is, if the time intervals is long relative to the velocity, theecho signal will change drastically precluding measurement of time shiftand decorrelation.

[0032] Echo decorrelation is mainly a function of the beamcharacteristics (the width of the beam among others). For example, for abeam with beam width of 1 mm, the echo signal would be totallydecorrelated after a 1 mm displacement of the scattering medium;however, a different transducer with a beam width of 2 mm would maintainsome of the correlation after a 1 mm displacement since the scatterersare still within the beam width.

[0033] Beam characteristics are range dependent. Consequently, anultrasound beam exhibits a range-dependent decorrelation characteristic.By experimentally or theoretically assessing the decorrelation for atransducer at all ranges and for displacements in all directions weobtain what we can call the “characteristic decorrelation” of the beam.Once the characteristic decorrelation has been assessed, measureddecorrelations in blood or tissue can be converted to displacement. Forexample, for a given range and direction of displacement across aparticular beam, a characteristic decorrelation curve is illustrated inFIG. 3; in this example, a measured decorrelation value of ρ₁ wouldcorrespond to a displacement of d=0.1 mm.

[0034] Thus, the decorrelation characteristic of the transducer servesthe purpose of a calibration factor which can be used to convertmeasured decorrelations into displacement and velocity.

[0035] One or more echo decorrelation values can be involved in thecomputation of a velocity value. Using a single decorrelation valuebetween a single pair of echo signals, the velocity is computed as theratio of the displacement obtained from the characteristic decorrelationat a given time interval. Using the example in FIG. 3, of the timeinterval between the echo acquisitions giving rise to the decorrelationof value ρ₁ was ΔT=1 ms, then the velocity v would be d/ΔT=0.01 mm/0.001s=10 cm/s. In practice the decorrelation characteristic versusdisplacement function may not be easily described analytically i.e., bya formula or the formula may not be suitable for inversion (obtaindisplacement from decorrelation). However, pre-calculated values ofcharacteristic decorrelation for small increments in displacement can bearranged in a “look-up” table: then, by “looking up” the decorrelationvalue in the table, the corresponding displacement can be obtained. Thisis termed the look-up-table (LUT) method. The LUT method makes noassumption regarding the shape of the characteristic decorrelation.

[0036] Estimation of velocity based on a single decorrelation value isvery sensitive to errors caused by other sources of decorrelation thatare not related to motion. However, by using the rate of change from twoor more echo decorrelation values, improved accuracy and precision ofestimates of velocity may be possible. Although the improvement inprecision can be expected since there is an implicit averaging proceduretaking place, the improved accuracy is specific to this application.This particular scheme for averaging multiple decorrelation estimates ismentioned as an example and is not intended to exclude the possibilityto use other averaging procedures with similar improvement.

[0037] Curve-fitting algorithms can be applied where a curve is fittedover a number of decorrelation measurement: the simplest curve being astraight line.

[0038] The linear fit method is illustrated in FIG. 4, where

[0039] a) a straight line is fitted over two or more decorrelationvalues. Since, by definition the decorrelation must be 1 for zerodisplacement, a single parameter (i.e., the slope of the line, termedthe decorrelation slope), defines the curve formed by decorrelationmeasurements;

[0040] b) similarly, another straight line fit is performed in thecorresponding area of the characteristic decorrelation curve; thisyields a characteristic decorrelation rate; and

[0041] c) the ratio of the decorrelation slope and the characteristicdecorrelation rate yields the velocity of blood (in units of mm/s).

[0042]FIG. 4 shows:

[0043] (a) Beam decorrelation estimated by experiment or theory; hereshown only for one distance from the transducer.

[0044] (b) Decorrelation measured with ultrasound at five intervals oftime (dots) with straight line fit.

[0045] (c) Velocity estimate using the decorrelation rate for thetransducer from (a) and the decorrelation slope from the measurement (b)at corresponding depths.

[0046] When the curve formed by subsequent decorrelation measurement isnot well approximated by a straight line, the linear-fit approach canlead to biased estimation. However, it follows from the abovedescription that when a linear fit is inappropriate, a higher-order fitcan be applied and more than one parameter is required to describe thebest fit.

[0047] The LUT method with multiple decorrelation estimates involves thesteps of

[0048] a) obtain from the characteristic decorrelation LUT thedisplacement for each measured decorrelation value obtained atsubsequent time intervals; and

[0049] b) a straight line fit is performed on the displacement versustime interval plot. The slope of the straight line fit is the velocityof blood.

[0050] An alternative approach to that of fitting of a particular trendto the decorrelation measurements is to simply calculate the average ofall available decorrelation estimates. It is important to recollect thatthe basic unit of the correlation algorithm is the cross-product of twoecho signals. For a pair of discrete echo signals s₁(i) and s₂(i), thecorrelation coefficient is given by$\rho = {\frac{\sum\limits_{i}{s_{1} \times s_{2}}}{\sqrt{\sum\limits_{i}{s_{1}^{2}{\sum\limits_{i}s_{2}^{2}}}}}.}$

[0051] Multiple decorrelation estimates can be made simultaneously bysquaring the sum of a number of echo signals, since such operationyields a sum of cross-products. In order to illustrate this, weestablish the relationship between the correlation coefficient and thenormalized sum of squares as follows:${S = \frac{\sum\limits_{i}\left( {s_{1} \times s_{2}} \right)^{2}}{{\sum\limits_{i}s_{1}^{1}} + {\sum\limits_{i}s_{2}^{1}}}},\quad {and}$${S = {\frac{{\sum\limits_{i}s_{1}^{2}} + {\sum\limits_{i}s_{2}^{2}} + {2{\sum\limits_{i}{s_{1}s_{2}}}}}{{\sum\limits_{i}s_{1}^{2}} + {\sum\limits_{i}s_{2}^{2}}} = {1 + \frac{2{\sum\limits_{i}{s_{1}s_{2}}}}{{\sum\limits_{i}s_{1}^{2}} + {\sum\limits_{i}s_{2}^{2}}}}}},$

[0052] When the two terms in the denominator of the above equation aresimilar, their arithmetic and geometric mean values can also be assumedto be similar, thus${1/{2\left\lbrack {{\sum\limits_{i}s_{1}^{2}} + {\sum\limits_{i}s_{2}^{2}}} \right\rbrack}} \approx {\sqrt{\sum\limits_{i}{s_{1}^{2}{\sum\limits_{i}s_{2}^{2}}}}.}$

[0053] Then we can write

S≈1+ρ.

[0054] With the above formula, decorrelation can be estimated from thenormalized sum of squares. When the squared sum involves more than twoterms, additional factors weighting the cross-products arise and dependon the number of terms in the square sum. For example, for the squaresum of three echo signals obtained at intervals of time Δt, we get${\sum\limits_{i}\left( {s_{1} + s_{2} + s_{3}} \right)^{2}} = {{\sum\limits_{i}\left\lbrack {s_{1}^{2} + s_{2}^{2} + s_{3}^{2} + {2s_{1}s_{2}} + {2s_{2}s_{3}} + {2s_{1}s_{3}}} \right\rbrack} \approx {\sum\limits_{i}\left\lbrack {s_{1}^{2} + s_{2}^{2} + s_{3}^{2} + {4s_{1}s_{2}} + {2s_{1}s_{3}}} \right\rbrack}}$

[0055] since we can assume that${\sum\limits_{i}{s_{1}s_{2}}} \approx {\sum\limits_{i}{s_{2}{s_{3}.}}}$

[0056] Thus the correlation for echo signals spaced by a single timeinterval has a weighting factor of four, while for a spacing of two timeintervals the weighting factor is two. This differential weighting ofcorrelation at different time intervals must be taken into account toobtain an accurate estimate of the average decorrelation.

[0057] Similarly, decorrelation estimates can be obtained from thesquared difference between echo signal pairs:${D = \frac{\sum\limits_{i}\left( {s_{1} - s_{2}} \right)^{2}}{{\sum\limits_{i}s_{1}^{2}} + {\sum\limits_{i}s_{2}^{2}}}},$

[0058] and following similar steps as above, we obtain

D≈1−ρ

[0059] With the above formula, decorrelation can be estimated from thenormalized squared difference of echo signals. In order to furthersimplify the calculation requirements, it is possible to substitute theabsolute difference for the squared difference. In this way, theoperation of calculating square values is avoided.

[0060] Note that while prior art teaches that decorrelation can be usedto assess displacement, the improvement in accuracy due to thecombination of multiple decorrelation values is novel.

[0061] For the purpose of measuring volumetric flow, the velocitycomponent of the flow normal (perpendicular) to the scan plane must beassessed. Alternatively, the angle between the scan plane and the flowmust be assessed.

[0062] This is illustrated for the vascular application in FIG. 5. Theblood velocity must be computed for blood velocity imaging, but thenormal flow component must be computed for volumetric flow estimation.

[0063] In general, the direction of flow can have any arbitrary anglewith respect to the reference axes of the ultrasound beam. This isillustrated in FIG. 6 where the direction of flow is shown relative tothe three-dimensional orthogonal axes centered on the transduceraperture. These axes in combination with the direction of the scan(sweeping of the beam) give rise to three spatial directions that aretermed axial, lateral, and elevational (FIG. 6). When a scatteringmedium displaces exclusively along the axis or the transducer (axialdisplacement), the motion can be assessed from the time shift of theecho signal when axial decorrelation is sufficiently low. When themotion occurs exclusively across the axial direction of the transducer,the echo signal decorrelates as the original scattering blood particlesmove out of the beam and new blood particles move into the beam.Displacement across the beam and in the scan plane is termed lateraldisplacement. Displacement across the beam and across the scan plane istermed elevational displacement.

[0064] Axial, lateral and elevational displacements introduce echodecorrelation giving rise to axial, lateral and elevationaldecorrelation components. Additionally, axial displacement introduces atime shift. Thus, in order to fully characterize displacement withrespect to a scan plane, all these decorrelation components must becomputed. Axial and elevational displacement can be measured asdescribed above. Lateral displacement can be estimated by“cross-decorrelation”, that is decorrelation analysis between-echosignals from adjacent beam locations within the scan plane. Therelationship between these decorrelations must also be assessedexperimentally or theoretically.

[0065] In a hypothetical example, assuming a linear decorrelation withdisplacement, lateral and elevational decorrelations can be combined inthe squared sense (that is, the square of the total decorrelation is thesum of the square of the lateral and elevational decorrelations); then,knowing the lateral displacement from cross-decorrelation analysis theelevational component of displacement can be isolated.

[0066] Analogously, the presence of axial, lateral and elevationalvelocity components can be considered. It is important to note that forthe purpose of measuring volumetric flow, the velocity component normalto the scan plane must be assessed. When the direction of flow is notperpendicular to the scan plane (elevation direction), flow velocityestimated with decorrelation may lead to biased estimation of volumeflow (unless the angle between the direction of flow and the scan planeis assessed).

[0067] In the common practice of intravascular ultrasound, the maincomponent of displacement is in the elevational direction. Thus, thecontribution of axial and lateral displacements may be neglected in somecircumstances without significant deterioration of the velocityestimates However, the contribution of lateral and axial componentsshould be kept under consideration. These components can increase duringthe examination of curved vessels and when secondary flow is present,among other possibilities.

[0068] Echo signal decorrelations also occurs due to sources that areunrelated to motion. For example, electronic noise present in the echosignals results in decorrelation. Two echo signals can differ only dueto corruption of independent realisations of the noise source. Thus, forimproved estimation of velocity from decorrelation, decorrelation fromsources of non-motion-related decorrelation must be isolated.

[0069] By computing the decorrelation of echoes from stationary tissue(e.g., vessel wall), the decorrelation due to sources other that motioncan be assessed. Then, this decorrelation can be deducted from the totaldecorrelation measured from moving blood (under the assumption thatblood and tissue echo signals contain the same amounts of electronic andquantization noise).

[0070] The main goal of the invention is to determine volumetric flowand velocity of blood with respect to the vessel through which it flows.However, the above scheme estimates velocity with respect to the beamunder the assumption that the transducer is fixed in position within thevessel. Normally, the transducer can and will move with the pulsation offlowing blood, adding an undesired motion component. In order to improvevelocity estimation, the motion of the transducer with respect to thevessel wall must also be assessed. Transducer motion can be assessed bymeasuring the time shift of echoes from quasi-static (relative to thehigh velocity blood) vessel wall tissue.

[0071] In general, the local velocity of blood varies within the vessellumen; that is, blood tends to move slower near static features (e.g.,the vessel wall, the ultrasound catheter) and faster away from thosefeatures (e.g., towards the center of the free lumen). Thus, to fullycharacterize flow it is important to estimate blood velocity in alocalized manner (note: commonly used Doppler-based techniques for flowestimation are based on the estimation of the peak value of the velocitywithin the lumen and the assumption of a velocity distribution).

[0072] Blood motion measurement scan be achieved in a localized mannerby analysis of gated (i.e., for “small” segments of) echo signals whichcorresponds to the local motion of small parts of the blood volume Forexample, FIG. 2 shows gated signals representing echoes from 0.25 mm ofblood. Many such local estimates of blood are measured in adjacentregions along the beam and can be combined to form a “velocity profile”.Velocity profiles obtained at different flow velocities are shown inFIG. 7: note the higher velocities near the center and low velocitiestowards the static wall. Adjacent beam positions in a scan plane can beused to form a map or image velocity. The sweeping of the beam(scanning) can be achieved.

[0073] a) by physically moving the transducer,

[0074] b) by physically moving a mirror which reflects the beam of astatic transducer, or

[0075] c) by electronically generating of the beam using an array oftransducers elements.

[0076] Velocity images can be superimposed on the gray-scale image ofstatic tissues by coloring image pixels according to the magnitude offlow. For example, from an experiment conducted in a live pig, four flowvelocity images computed within one heart cycle are shown in FIG. 8a.Thus, unlike any previously described method, an image of the velocityof blood flowing normal to the cross-section of the scan plane isobtained (illustration in FIG. 1, measurement in FIG. 8).

[0077] Additionally, velocity components or the volume flow can beconverted into an audible signal in order to provide a different way topresent flow information. For example, a converter could be used totransform flow data into an analog signal which, followingamplification, could drive a loudspeaker. This may be a useful featurewhen the operator is unable to look at the monitor while manipulatingthe ultrasound transducer. For instance, the range and mean value flowvelocity could be represented by the bandwidth and the pitch of theoutput sound. This presentation of flow information is analogous to whatis used in Doppler systems where the frequency of the Doppler signal isby default in the audible range and needs only be amplified andconnected to a speaker. However, here the sound is synthesized fromindependently computed flow information.

[0078] Additionally, decorrelation based estimation of motion can beused to differentiate areas where there is moving blood (that is, thefree lumen area) from areas where tissue is static. Blood flows at amuch higher velocity relative to the motion of healthy or diseasedtissue (such as the vessel wall, plaques and dissections).

[0079] The prior art in the intravascular ultrasound applicationdescribes the combined use of two catheters to measure volumetric flow.A first intravascular ultrasound catheter is used to measure free lumenarea and is not capable of measuring flow. A second Doppler-ultrasoundcatheter is used to assess the velocity of blood flow and is not capableof measuring the area of flow. Additionally, the Doppler catheter cannotmeasure local velocity at many spatial locations: a velocity profilewithin the lumen is not measured but assumed based on a singlemeasurement of velocity (peak or mean).

[0080] The combination of measured area and assumed velocitydistribution yields volumetric flow. However, the cross-sectional areameasurement and the flow estimation are performed at two spatiallyseparated locations. Alternatively, cross-sectional area can be measuredin one location of the vessel, and later the velocity can be measured inthat same location. Thus, the prior art is limited to either“simultaneous” or “in place” measurements, but does not teach both.

[0081] In the current invention integration of the measured velocity mapover all points in a cross-section of the free lumen area yields thevolume of blood flowing through the vessel. Since the area ofintegration and the flow velocity is computed from the same signals, itis a fact that the estimation of area and velocity are simultaneous andin place.

[0082] Alternatively, a reduced, but representative number of pointswithin a cross-section of the free lumen area can be used to estimatethe average value of flow velocity in the entire free lumen area. Theaverage velocity can be estimated from a partial area of the flow areawhen the velocity distribution is known to first approximation.Simultaneously, the area of flow can be calculated at the samecross-section. The area of flow times the average flow yields the volumeof blood flowing through the vessel. Like in the Doppler approach, avelocity profile within the lumen must be assumed to calculate theaverage velocity based on a restricted number of measurements ofvelocity within the lumen. However, unlike the Doppler approach, in thisalternative implementation velocity and flow are measured simultaneouslyand in place.

[0083] Since this measurement can be performed at regular intervals thatare small compared to the period of the heart or respiratory cycle,phasic volumetric flow can be assessed (phasic meaning the time historywithin a cycle). Phasic volumetric flow measured in a live-pigexperiment is shown in FIG. 8b.

[0084] The relationship between local blood pressure and the volumetricflow yields additional hemodynamic information of extreme clinicalrelevance. Particularly, the change of the pressure-flow relationship inresponse to vaso-active drugs is used to investigate the reaction ofdifferent parts of the vasculature. For this reason, several methodshave been developed that combine pressure-sensing catheters withintravascular Doppler-ultrasound assessment of volumetric flow.Alternatively, lumen area or diameter can be used as an estimator ofpressure and velocity can be used as a measure of volumetric flow. Alimitation common to current multi-variable methods is the inability toassess the hemodynamic variables coincidentally in time and space.Additionally, multiple catheter approaches suffer from possibleinterference between devices: for example, in the combination ofintravascular ultrasound for area measurement and Doppler catheter forvelocity measurement, the intravascular catheter disturbs the blood flowand therefore affects the velocity that is measured by the Dopplercatheter.

[0085] Since the endoluminal pressure is intimately related to thechange in cross-sectional lumen area (particularly, there is a linearrelationship between pressure and lumen diameter toward late diastole),this invention can provide concurrent and coplanar measurements ofcross-sectional lumen area and volumetric flow. This is shown in FIG. 9where phasic cross-sectional area and phasic volumetric flow werecalculated simultaneously and from the same scan plane from the live pigexperiment. The phasic relationship between area and flow can provideinformation on the resistance of the vasculature. Spectral analysis ofphasic area and flow can be used to assess arterial impedance.

[0086] So far the measurement of flow has been described assuming thatthe backscatter from blood is uniform within the lumen and in time.However, normally red blood cells tend to form clusters, a processcalled aggregation, and arrange themselves in “strings” called rouleaux.In the following blood, rouleaux are positioned along the direction offlow. The presence of RBC aggregation and rouleaux is a function of thecyclic variation of the local shear: regions of high shear (near staticstructures) have low aggregation and areas of low shear (near the centerof the free lumen) have high aggregation. Thus, a spatial distributionof RBC aggregation as well as a cyclic temporal variation with the heartrate are known to exist. In intravascular ultrasound imaging, theseeffects are manifested as a cyclic and spatial variation of the echointensity (since larger aggregates of RBCs also backscatter strongerechoes). Thus, examination of the echo intensity and the backscattercoefficient function can yield information about the scattereraggregation.

[0087] The shape and size of the back scattering particles can also havea significant effect on the decorrelation phenomenon. Clearly, the echofrom a single point scatterer moving across a beam will decorrelatefaster than the echo arising from a long string of aligned scatterersmoving across the same beam. Thus, the dependence of decorrelation onblood backscatter should be compensated for improved accuracy in volumeflow and velocity estimation. For example, long rouleaux or clusterspresent in the central (low shear) part of the free lumen would resultin local underestimation of velocity.

[0088] Analogously to the determination of the characteristicdecorrelation of an ultrasound beam, for each shape and size ofscatterers we can obtain a “scatterer characteristic decorrelation”function. Then, from the scatterer type/shape can be estimated frombackscatter analysis and the velocity estimates can be compensated forscatterer characteristic decorrelation.

[0089] The effect of RBC aggregation on decorrelation is a function ofthe size of the aggregate relative to the wavelength. Therefore,applications that utilize high frequency ultrasound (i.e., wavelengthssimilar to the aggregate size) may be expected to experience a higherdependence on aggregation. Conversely, low ultrasonic frequencies (i.e.,wavelengths much larger than the RBC diameter) may be expected to beless affected by aggregation-dependent decorrelation. In practice,reasonable estimates of flow can be obtained withoutbackscatter-dependent compensation.

1. A method to measure spatial fluid flow components and their velocityprofiles in a number of locations in a cross-sectional area of a lumenor other body cavity by using ultrasound characterized in that: a) thecross-sectional area is interrogated by a plurality of ultrasonic beams;b) the estimation of the spatial flow components is obtained from acombination of estimations of axial, lateral and total flow; c) theestimation of one or more flow components is obtained through time-shiftand decorrelation analysis of two or more beam-signals of theinterrogating ultrasound transducer; and d) the estimation accuracy isfurther improved by the use of a reference decorrelation curve obtainedfrom experiments or beam-theory or both.
 2. The method of claim 1,further characterized by simultaneous and in place estimation by thesame means of the dimensions (cross-sectional area, mean diameter, etc.)of the flow area.
 3. The method of claim 2, further characterized bycombining the component of flow velocity normal to a cross-section of anarea of flow in a number of spatial locations and the flow area asdefined in claim 2 to derive a momentary value of the volume flow. 4.The method of claim 2, further characterized by combining (1) theaverage over a number of spatial locations of the component of flowvelocity normal to a cross-section of an area of flow and (2) the flowarea as defined in claim 2 to derive a momentary value of the volumeflow.
 5. The method of claim 2, further characterized by combining thecomponent of flow velocity normal to a cross-section in a number ofdiscrete spatial locations and the corresponding areas of said discretespatial locations to derive a momentary value of the volume flow.
 6. Themethod of claim 5, wherein the step of combining normal velocity andcorresponding discrete areas is an integration operation.
 7. The methodof claim 4, further characterized by calculation of the averagecomponent of flow velocity normal to a cross-section of a total area offlow from a partial section of said total area of flow.
 8. The method ofclaim 7 wherein said partial section of said area of Flow is a singleline segment of said total flow area.
 9. The method of claim 3-8,wherein a time history of volume flow and lumen dimensions are recordedwithin the heart cycle or the respiratory cycle.
 10. The method of anyof claims 1 to 9 characterized by the use of means to compensate for thedecorrelation characteristic of the scattering medium.
 11. The method ofany of claims 1 to 10 characterized by the use of means to compensatefor non-motion-related decorrelation components.
 12. The method of anyof claims 1 to 11 characterized by the use of means to compensate thevelocity measurements for relative motion of the transducer with respectto relatively static tissues.
 13. The method of any of the precedingclaims wherein the velocity is measured from the slope of thedecorrelation.
 14. The method of any of the preceding claims wherein thevelocity is measured from the slope of decorrelation-deriveddisplacements.
 15. The method of any of the preceding claims wherein thedecorrelation is derived from squared differences of echo signalcalculations.
 16. The method of any of the preceding claims wherein theaverage decorrelation is derived from squared sums of echo signals. 17.The method of any of the preceding claims wherein the velocity, velocityprofile or volume flow measurements are converted into an audible sound.18. The method of any of the preceding claims wherein the transducer isan intraluminal ultrasound transducer where the beam is scanned in aplane by one of the following a) rotating a transducer; b) rotating amirror which redirects the beam of a static transducer; c)electronically generating the beam through a multi-element transducerwithout transducer motion; and d) a combination of a) and c).
 19. Themethod of any of the preceding claims wherein the ultrasound beam isscanned through adjacent spatial locations in a sequential step-wiseway.
 20. The method of any of the preceding claims wherein theultrasound beam is scanned through adjacent spatial locations in anon-sequential way.
 21. The method of any of the preceding claimswherein the ultrasound beam is scanned through adjacent spatiallocations in a back and forth way.
 22. The method of any of thepreceding claims wherein the transducer is a single element, a linear,annular or phased array.